This invention is directed in principal part to an ultrasonic transducer device. More particularly, this invention is directed to a dual mode ultrasonic transducer device, one that can be used for both diagnostic investigations and therapy.
Ultrasound is widely used in modern medicine for diagnostics and treatment in such fields as obstetrics, cardiology, endocrinology, gastroenterology, neurology, ophthalmology, urology, osteoporosis, and clinical diagnostics. A wide range of clinical trials are being conducted for breast tumor ablation, urine fibroids ablation, benign prostatic hyperplasia ablation, fibrillation, cardiac, bleeding control, and brain disorder treatments (Clement, 2004, Perspectives in clinical uses of high intensity focused ultrasound, Ultrasonics, 42, 1087-1093).
Ultrasound diagnostics uses low-power ultrasonic scanners for investigation and visualization of inner organs, layers and structures, for determination of blood flow direction and velocity, for measurement of density and other parameters of tissues, and for detection of cancer and other tumors. Following an ALARA (as low as reasonably achievable) principle, diagnostic evaluation spatial peak temporal average intensities do not exceed 100 mW/cm2 (Kremkau 2006, Diagnostic ultrasound: principles and instruments, 7th ed. Philadelphia Pa., Saunders).
Ultrasound propagating through the tissues attenuates and creates heat proportional to ultrasound intensity. The high intensity focused ultrasound (HIFU) can kill tissue through coagulative necrosis. This idea of HIFU for tissue necrosis and therapy actually predates its suggested use as a diagnostic tool. Nevertheless, it is only recently that therapeutic HIFU procedures have become practical and reliable, predominantly due to significant advances in ultrasound imaging technology over past decade that have enabled near real-time non-invasive monitoring and control of ultrasound treatment. At the same time, ultrasound guidance became a viable and low cost alternative to magnetic resonance imaging (MRI), X-ray computer tomography (CT). Ultrasound offers a credible potential to control the HIFU ablation process, and a number of ultrasound thermal imaging methods are currently under way to quantify HIFU induced temperature changes to tissue properties (e.g. Hill and Ter Haar, 1995, Review article: HIFU potential for cancer treatment. Br. J. Radiol, 68, 1296-1303; Zheng and Vaezy, 2010, An acoustic backscatter-based method for localization of lesions induced by HIFU, Ultrasound in Med & Biol, 36, 4, 610-622). Consequently, medical ultrasound imaging establishes itself as a vital component of the HIFU therapy, and gains clinical acceptance for a safe and effective tissue ablation and cancer therapy (Haar, 2010, Ultrasound bioeffects and safety, Proceedings of the Institution of Mechanical Engineers, Part H: Journal of Engineering in Medicine, 224 (2), pp. 363-373).
Being a subject of research for many year, ultrasound therapy only recently has found a widespread use in medical applications (ter Haar G, 2007, Therapeutic applications of ultrasound, Progress in Biophysics and Molecular Biology 93, 1-3, 2007, 111-129). Ultrasound therapy uses considerably higher ultrasound intensities than imaging. Levels up to 10 W/cm2 are used for fast overheating of local areas of tissue. In hyperthermia treatment of cancer and tumors, the tissue is heated using ultrasound to temperatures of 43-45° C. for several minutes. Under such conditions tumor cells become much more susceptible to radiotherapy and chemotherapy. In physiotherapy ultrasound is used to increase the elasticity of sinews and scars, improve the mobility of joints, provide analgesic effects, alter blood flow, and produce muscular spasms. High intensity ultrasound (10-2000 W/cm2) is used for tissue cutting, thermal ablation and for arresting internal bleeding (hemostasis) due to blood fibrillation. Piezoelectric and magnetostrictive transducers are widely used to transform an electrical current and voltage into mechanical oscillations that generate an ultrasound field.
Therapeutic ultrasound targets deep tissue using focused transducers with one or more active elements typically emitting continuous wave signals. Ultrasound can also be focused by manipulating the driving electrical signals (phase and amplitude) of multiple elements. (Cathignol, 2002, High Intensity Piezoelectric Sources for Medical Applications: Technical Aspects, Nonlinear Acoustics at the Beginning of the 21st Century, 1, 371-378.) Larger elements are more economical, but require mechanical steering and suffer a loss of acoustic efficiency due to the heating and presence of parasitic Lamb waves on their surfaces. (Kluiwstra et al., 1997, Design Strategies for Therapeutic Ultrasound Phased Arrays, SPIE International Medical Imaging Symposium). Properly energized and poled small elements are more economical but require complex circuitry. Resultant phased arrays comprising the small ultrasound elements can steer acoustic focus electronically, with most of the complexity, cost and associated quality assurance activity being shifted to assembly processes and driving system manufacturing.
It is to be noted that diagnostic imaging imposes a set of design requirements contradictory to a therapeutic mode of operation. While the continuous therapeutic mode favors narrow band, sharp resonant transducers for a high power output and improved efficiency over extended procedure time, the diagnostic mode relies predominantly on the pulse signals and favors broadband transducer design. A broader bandwidth results in more energy in short imaging pulses and translates into more sensitivity, yet broader bandwidth indicates less efficient vibration at resonance frequency. Because of these contrary requirements, spatially separated and individually operated therapeutic and imaging elements are used in conventional dual-mode HIFU applicators. Another imaging transducer design challenge is to match the acoustic impedance of about 33 MRyals, typical of PZT ceramics, to the relatively low 1.5 MRyals impedance of water or tissue in order to obtain a short impulse response and broad bandwidth. Typically, this impedance matching is accomplished by fabricating multi-stage close to quarter wave matching layers using epoxy resins loaded with tungsten or alumina powders. (Kosoff, 1966, The Effects of Backing and Matching on the Performance of Piezoelectric Ceramic Transducers, IEEE Transactions on Sonics and Ultrasonics, SU-13, 1, 20-30) Based on KLM transmission line models (Krimholtz, Leedom, Matthaie 1970, New Equivalent Circuits for Elementary Piezoelectric Transducers, Electron Lett, 6, 13, 398-399), the acoustic matching layers of imaging transducers should have impedances in the range of 8 to 15 MRyals. This is difficult but feasible to achieve using epoxy with sufficiently dense powder loading. Conversely, in the case of therapeutic transducers that operate near resonance to maximize surface vibration, the acoustic impedance of a matching layer should be much less than the impedance of the transducer material and water. Only few such practical low impedance materials available (Toda, 2002, Narrowband Impedance Matching Layer for High Efficiency Thickness Mode Ultrasonic Transducers, IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, 49, 3), favoring the designs of therapeutic transducer without matching layers at all.
For most medical imaging ultrasonic applications one requires the large pulse amplitude for effective transmission through a body of a patient and good signal to noise ratio of the detected signal. This is especially important since biological tissues are attenuative or scattering materials. Also, the short pulse duration is necessary for good axial resolution, i.e. two signals in a short distance should be detected without interference. Consequently, in imaging transducers it appears desirable to dampen the vibration of the piezoelectric transducer material as quickly as possible to prevent ringing and multiple reflections. On the other hand, in therapeutic applications liquid cooled undampened sharp resonance transducers are typically used to produce acoustic signals of sufficient magnitude over a length of time. When using high power resonance transducers for imaging, improved power transmission leads to an improved imaging sensitivity, while ringing limits axial resolution. Vice versa, dampened transducers typically have lower power transmission and sensitivity on account of higher axial resolution. This trade off between power transmission and axial resolution presents another challenge for a design of single element transducer for simultaneous therapeutic and imaging application.
Thereby, multiple design constraints such as acoustic matching of high impedance ceramic transducers for broadband imaging and narrow band therapy, are fundamentally different, making the design and fabrication of the dual mode transducers for imaging and therapy extremely challenging.
At present, the ultrasound based imaging and Magnetic Resonance Imaging (MRI) are the competing modalities that offer a potential to evaluate the location of thermally induced lesions in a patient body. Due to restrictive technological complexity of combining HIFU and MRI, the real time ultrasound-based detection of the high lesions is favored. Some early methods relied on the speed of sound change with temperature to control the HIFU exposure (Miller, Bamber, Meaney, 2002, Functional limitation of non-invasive temperature imaging by means of ultrasound echo strain estimation. Ultrasound Med & Biol, 28, 13 19-1333). However, the accuracy of this method is poor due to significant inter-patient variability, non-linear acoustic effects, thermal expansion, limited data available for different tissues. Other methods utilized the appearance of hyperechoic regions in B-mode imaging (Vaezy, Shi, Martin, Chi, Nelson, Bailey, Crum, 2001, Real-time visualization of HIFU treatment using ultrasound imaging, Ultrasound Med & Biol, 27, 33-42), which was only applicable when there was a significant cavitation.
Moreover, human body supports the propagation of many kinds of ultrasound waves, each of which can be used to acquire an image of internal organs. For example, compression waves reveal a tissue's density, while shear waves reveal tissue's elasticity. So called “harmonic imaging” has become common to ultrasonic medical diagnostics due to a higher resolution of the second harmonic in comparison with the fundamental frequency. The tissue response to harmonic vibration is characterized by Lamé parameters: λ, which is associated with the elastic resistance to volume change, and μ, which characterizes the tissue's elastic resistance to shape change. A typical value for shear modulus μ is on the order of kPa, while a typical value for λ is about 2.3 GPa, which is about six orders of magnitude higher than the shear modulus. In that limit, compression waves have a speed νP=√{square root over ((λ+4/3μ))}≈√{square root over (λ/ρ)}≈1500 m/s and can propagate in ultrasound (megahertz) frequency range, while shear waves are characterized by significantly lower speed νS=√{square root over (μ/ρ)} and propagate at low sonic (kilohertz) frequencies. Because the soft tissue is 70-80% water, its resistance to volume change, λ, and density, ρ, do not vary much in comparison to μ (Sarvazyan A P, Rudenko O V, Swanson S D, Fowlkes J B and Emelianov S Y, 1998, Shear wave elasticity imaging: a new ultrasonic technology of medical diagnostics. Ultrasound in Med. & Biol. 24 1419-1435). Consequently, shear waves can be a good tool for quantitative evaluation of deep-organs stiffness, providing an important palpatory diagnosis information. These, along with the aforementioned differences between imaging and therapeutic transducer design requirements, rapidly shifts a focus to the implementation of the nonlinear multiwave wave elastography in medical imaging and nondestructive testing (e.g. Fink and Tanter, 2010, Multiwave imaging and super resolution, Physics Today, 63, 2; Brysev et al, 2004, Nonlinear ultrasonic phase-conjugate beams and their application in ultrasonic imaging, Acoust. Phys. 50, 623-640; Mathias Fink. 2002, Acoustic Time-Reversal Mirrors. Topics Appl. Phys. 84, 17-43). Besides A-scan, B-scan, Doppler imaging, harmonic imaging, the recent advances also include contrast imaging, 3 and 4 dimensional imaging, coded excitation and elastography. The state of the art ultrasound systems use hundreds of piezoelectric transducers and ultrafast scanners (Sandrin L, Tanter M, Catheline S, Fink M. Shear modulus imaging with 2-D transient elastography, IEEE Trans Ultrason Ferroelectr Freq Control. 2002 April; 49(4):426-35.) to form a high-resolution image. Unlike conventional ultrasound imaging, where multiple bursts are required to produce an image, shear modulus imaging technique numerically reconstructs an image after each ultrasound transmission. The reconstruction process relies on the time reversal of the digitized compression back-scattered waveforms. A similar process has been described in a transient source time reversal acoustic holography (Sapozhnikov O A, Ponomarev A E and Smagin M A, Transient acoustic holography for reconstructing the particle velocity of the surface of an acoustic transducer, Acoustical Physics, 52, 3, 2006), which also showed that a quality of reconstruction depends on individual transducers' directivity. An array with ultrasound transducers having different directivities can substantially improve the accuracy of the time reversal imaging reconstruction or focusing. Transducers can be individually poled, either randomly or in an order sequence, and oriented in such a way that they sense a compressional wave coming from predefined direction. Alternatively, piezoelements can be made to sense both compressional and shear waves simultaneously. For example, 10° rotated Y-cut overtone-polished parallel-plated LiNbO3 crystals were used as dual-mode source and receiver transducers in ultrasound interferometric measurements (Sinelnikov Y D; Chen G R; Liebermann R C, Dual mode ultrasonic interferometry in multi-anvil high pressure apparatus using single-crystal olivine as the pressure standard, High Pressure Research, 24, 1, 2004, 183-191).
It is also very desirable that the imaging array transducers have broad bandwidth and high sensitivity, and it is predicated on the ability to eliminate mechanical movement and deflect the beam electronically. Typically, imaging transducers are used both for generating the pulses and detecting respective echoes that arise when ultrasound pulses are partially reflected at boundaries between structures with different characteristic impedances. Elements of a group in an array can be excited in succession so that the ultrasound beam is electronically moved across the face of the transducer, providing an image similar to that obtained by scanning a single element transducer manually. An ability to steer and focus the acoustic energy at one or more locations simultaneously by manipulating the phase and amplitude of each element makes multiple elements ultrasound array attractive for therapeutic applications.
However, the material properties impose a significant constraint on the design of ultrasound imaging and therapeutic arrays. The conversion efficiency of the transducer indicates how well the transducer converts both the applied voltage into ultrasonic pressure pulse and the received echo into electrical voltage. Related to conversion efficiency, transducers' sensitivity is defined as the product of transmit and receive efficiencies. The sensitivity of a single element is largely defined by the piezoelectric material constants d33 and g33, which are indices of how well the material converts the voltage signal into mechanical deformations and the mechanical stress back into electrical voltage, respectively. Typically it is not feasible for both d33 and g33 constants to be large and typically a combination of closely spaced specially selected piezomaterials can be beneficially used in transmit and receive mode. The use of such piezocomposite materials becomes important for ultrasound guided HIFU therapy that relies on ultrasound accurate spatial localization of HIFU-induced lesions in real time and after HIFU exposures. The main objective shifts from the visualization of static internal organs of to the monitoring of target ablation process and providing a timely operator feedback for the treatment planning.
On the other hand, the success of the HIFU procedures depends on the intensity gain of the transducer, while the optimum imaging sensitivity, penetration, and ability for a similar dynamic focusing need to be preserved. In most designs the intensity is maximized by increasing the total power, which is proportional to a surface of emitting transducer. However, large transducers suffer from increased vibration and scattering losses, while a combination of small transducers is difficult to handle. One solution to obtain the high therapeutic intensity gain at a target is to arrange the therapy and imaging elements in a non-contiguous way. In such interleaved arrangement all array therapy and imaging elements point in the same direction, making the dynamic scanning region essentially similar for both arrays thus improving localization and temperature monitoring of the HIFU application. Such interleaved dual mode array will enable an improved multimode imaging that relies on a generation of the sufficiently high intensity of acoustic beam by therapeutic elements, necessary for generation of harmonics on a path of the backward propagation to the imaging elements.